The field of the invention is nuclear magnetic resonance imaging (MRI) methods and systems. More particularly, the invention relates to MRI diffusion weighted imaging (DWI).
Any nucleus which possesses a magnetic moment attempts to align itself with the direction of the magnetic field in which it is located. In doing so, however, the nucleus precesses around this direction at a characteristic angular frequency (Larmor frequency) which is dependent on the strength of the magnetic field and on the properties of the specific nuclear species (the gyromagnetic constant gamma (of the nucleus). Nuclei which exhibit this phenomena are referred to herein as “spins”.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. A net longitudinal magnetization M0 is produced in the direction of the polarizing field, but the randomly oriented magnetic components in the perpendicular, or transverse, plane (x-y plane) cancel one another. If, however, the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net longitudinal magnetization, M0, may be rotated, or “tipped” into the x-y plane to produce a net transverse magnetic moment Mt, which is rotating, or spinning, in the x-y plane at the Larmor frequency. The practical value of this phenomenon resides in the signal which is emitted by the excited spins after the excitation signal B1 is terminated. There are a wide variety of measurement sequences in which this nuclear magnetic resonance (“NMR”) phenomena is exploited.
When utilizing NMR to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region which is to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles which vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the “MR” image using one of many well known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing magnetic fields (Gx, Gy, and Gz) which have the same direction as the polarizing field B0, but which have a gradient along the respective x, y and z axes. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified. MR imaging is employed to image a number of anatomical and physiological features of living animals.
Diffusion-weighted imaging (DWI) is a powerful MRI technique for probing microscopic tissue structure. In DWI, a pulse sequence is employed which contains a magnetic field gradient known as a diffusion gradient that sensitizes the MR signal to spin motion. A DWI pulse sequence is shown in FIG. 2. It includes the generation of a 90° selective RF excitation pulse 100 which is produced in the presence of a slice select gradient pulse 102 to excite spins in a 2D slice. A 180° RF refocusing pulse 104 is produced in the presence of a slice select gradient pulse 106 to refocus the transverse magnetization and an NMR signal 108 is acquired at that time in the presence of a readout gradient pulse 110. The pulse sequence is repeated and a phase encoding gradient pulse 112 produced just prior to signal readout is stepped through different values to sample k-space in the conventional manner.
Diffusion weighting of the acquired NMR signal 108 is provided by diffusion gradients 114, 116 and 118 applied along the respective slice select, phase encode and readout gradient axes. Each diffusion weighted gradient 114, 116 and 118 includes a first gradient lobe produced prior to the 180° RF pulse 104, and a second gradient lobe produced after the 180° RF pulse 104. The first and second diffusion gradient lobes are of equal size (area), but their relative values are changed to measure diffusion in different directions. In this spin echo sequence the two lobes of the diffusion gradient waveform are produced before and after the 180° RF pulse 104 and they have the same polarity. If the diffusion gradient is employed in a pulse sequence that does to have a 180° RF pulse, its waveform takes the form shown in FIG. 4. In this alternative embodiment the waveform is the same as that described above except the two lobes have opposite polarity.
In a DWI pulse sequence the detected MR signal intensity decreases with the speed of water diffusion in a given volume of tissue. The first moment of the diffusion gradient, also known as the “b-value” determines the speed of diffusion to which the image is sensitive. This b-value may be adjusted by either varying the area of the two lobes of the diffusion magnetic field gradient, or by varying the time interval between them. Referring to FIGS. 3 and 4, when water motion in the subject is unrestricted, the MR signal intensity at the center of the echo using a spin-echo diffusion-weighted pulse sequence is related to the b-value as follows:
                    A        =                                            S              ⁡                              (                b                )                                                    S              0                                =                      e                          -              bD                                                          (        1        )            where the “b-value” b=γ2G2δ2 (Δ−δ/3). The parameter γ is the gyromagnetic ratio and G is the amplitude of the applied diffusion magnetic field gradients. S(b) is the MR signal magnitude with diffusion weighting b, and S0 is the MR signal magnitude with no diffusion weighting (b=0). The parameter D is the diffusion coefficient of water within itself (in mm2/s), which directly reflects the fluid viscosity where there are no structural restrictions to diffusion of the water. As shown in FIGS. 3 and 4, Δ is the time interval between the onsets of the two diffusion gradient lobes and δ is the duration of each gradient lobe. The diffusion coefficient D in equation (1) may be calculated, since b is known and the attenuation A can be measured.
There are a large number of clinically scientifically important applications for DWI. These include early detection and characterization of cytotoxic edema caused by cerebral infarction, improved tumor characterization through detection of restricted diffusion within a cellular tumor, and cerebral “tractography” for fiber angle mapping of the cerebral white matter, as well as many others. Within the abdomen, low b-value DWI is commonly used for liver imaging, to null the signal from flowing blood to improve the conspicuity of liver lesions such as metastases or primary liver tumors.
Unfortunately, DWI is exquisitely sensitive to motion. Large phase shifts from small patient bulk translations are encoded by the large amplitude diffusion weighted gradients. These phase shifts are extremely problematic for most DWI imaging methods such as that described above and illustrated in FIG. 2, which are “multi-shot” methods, i.e., require multiple excitations and acquisitions in order to form a complete k-space matrix prior to Fourier transformation. The slightest amount of motion leads to large phase misregistrations between data acquired in the different shots due to first and higher order moment phase shifts from velocity and acceleration during the time that the diffusion pulses are being played. Such phase shifts lead to devastating image artifacts in the form of severe ghosting and blurring. For this reason, most DWI methods rely on echo planar imaging (EPI) which is a high SNR, ultra-rapid method that can acquire all lines of k-space in a single shot, avoiding problems of phase misregistration. However, EPI suffers from poor spatial resolution, severe distortion in areas of high magnetic susceptibility, ghosting artifacts, and requires high performance hardware. Overall, the poor image quality of EPI has greatly restricted the clinical use of DWI.